The invention relates to stents made of a material with an elongation at rupture of 30% or less and with a tubular base body which entirely or in parts comprises structural segments which are interconnected in the logitudinal direction of the stents by means of transverse connectors and in which stents the structural segments comprise a zigzagging or undulating structure of a brace which is wrapped around the longitudinal axis of the stent.
Coronary heart disease, in particular acute myocardial infarction, is one of the most frequent causes of death in Western Europe and North America. In more than 80% of cases, myocardial infarction is caused by thrombotic obstruction of a coronary artery as a result of rupture of atheromatous plaque in pre-existing stenosing atheromatosis. The following are decisive factors in the long-term prognosis after acute myocardial infarction:                an effective and long-lasting reopening of the infarction artery;        the duration of the thrombotic obstruction of the vessel;        prevention of a major myocardial insufficiency and ventricular remodelling;        gaining control over rhythmogenic complications.The above-mentioned factors not only determine cardiovascular mortality but also the quality of life following the infarction.        
Non-operative methods of stenosis treatment, in which among other things by means of balloon dilatation (percutaneous transluminal coronary angioplasty, PTCA) the restricted or obstructed blood vessel is opened up, have been established for more than twenty years. This procedure has proven reliable, in particular in the therapy relating to acute myocardial infarction. After the blood vessel has been expanded, in about a third of cases above-average proliferation occurs as a result of cell growth triggered by treatment, which proliferation finally leads to renewed angiostenosis (restenosis). The elasticity of the dilated blood vessel is a further cause of restenosis. After the balloon has been removed, the blood vessel contracts excessively so that the cross section of the vessel is reduced (obstruction, so-called negative remodelling). The latter effect can only be prevented, or at least impeded, by the placement of a stent.
In interventional therapy of stable angina pectoris with coronary heart disease, the introduction of stents has brought about a clear reduction in the rate of restenoses and thus to improved long-term results. This applies not only to primary stenosis but also to recidivation stenosis. An increase in the primary lumen gain is the major benefit of a stent implantation.
While the use of a stent can result in a more optimal vascular cross section, the presence of such a foreign object does however initiate a cascade of microbiological processes which can lead to a gradual closing up of the stent.
As a rule, a stent comprises a tubular base body with a circumferential wall made of braces and transverse connectors between which braces and connectors cells of various different shapes extend. A common feature of all typically used stents consists of the need, irrespective of the design of the stent, to allow expansion from a closed (non-expanded) state to an opened-up (expanded) state. In the non-expanded state, the stent is guided into the region of the lesion of the blood vessel. At the location of application the structure expands, either as a result of externally applied forces (e.g. by means of a balloon arranged in the interior of the stent) or as a result of the stent being designed so as to be self-expanding (e.g. by using a memory material).
A stent design must in particular meet the following requirements:                in its expanded state, a stent must evenly support as large a surface of the wall of the blood vessel as possible;        the design must support expansion of the stent in radial direction without allowing elongation in axial direction;        in the non-expanded and expanded state and during the transition from the non-expanded to the expanded state, the individual braces and transverse connectors should, as far as possible, be arranged in a common radial circumferential plane, i.e. any projection of individual structural elements must be suppressed so as to prevent injury to the blood vessel;        any failure of the material for the duration of the healing- and implantation period must be avoided.        
Modern stent designs take account of these requirements with the use of typically employed stent materials.
Initially, stents were predominantly made from surgical steel, e.g. 316L. Over time it became clear, however, that while the materials used were biocompatible, in the medium to long term some of them encouraged thrombosis formation while others encouraged adhesion of biomolecules to their surfaces. Stents made from a biodegradable material provide a starting point to solving these problems. The term “biodegradation” relates to hydrolytic, enzymatic and other metabolic decomposition processes in the living organism which lead to gradual decomposition at least of large parts of the implant. For example, a multitude of plastic materials have been proposed as stent materials, which, while providing good biocompatibility and good degradation behaviour, due to their mechanical properties have at best been limited in their use for medical applications.
To overcome this disadvantage, the use of special biodegradable metal alloys has been envisaged, as described in particular in U.S. Pat. App. Pub. No. 2002/0004060 and DE 199 45 049 whose disclosure is herewith fully incorporated in this patent specification. The metal alloys comprise special iron alloys, tungsten alloys and magnesium alloys. However, these favoured materials are at least partly associated with a disadvantage in that their elongation at rupture is 30% or less so that all typically applied stent designs are thus excluded. For example, material 316 L has an elongation at rupture of 40-50%, determined on a tubular tensile test sample with a diameter of 1.6 mm and a wall thickness of 0.1 mm, at a sample length of 60 mm and a measuring length of 22 mm. A typical biodegradable magnesium alloy has an elongation at rupture of 14%, determined on a tubular tensile test sample with a diameter of 1.6 mm and a wall thickness of 0.2 mm, at a sample length of 50 mm and a measuring length of 30 mm.
In this context the term “rupture” refers to separation, caused by overloading, of the binding between atomic components of a material along a surface of fracture which extends across the entire width of the material sample. At the time of rupture, the tensions built up in the material as a result of external or internal forces exceed the material-specific stress at failure and thus exceed the binding forces that are effective in the surface of the fracture. The resistance to fracture states the force needed for a material sample to be loadable to fracture. Fractures without deformation or with little deformation can be compensated for without any problems with the use of the above-mentioned metal alloys. In contrast to this, deformation fractures such as transverse ruptures and torsion fractures are unavoidable with conventional stent designs because the elongation at rupture of the material is insufficient.
The term “elongation at rupture” refers to the remaining extension after a tensile test up to rupture, i.e. the difference in the length between the original sample and that of the fractured sample.
It is an aspect of the present invention to provide a stent design which is suitable for materials with an elongation at rupture of less than 30%.